Evaluation of Canine Fracture Fixation Bone Plates

Transcription

1 Evaluation of Canine Fracture Fixation Bone Plates by Edward Kalust Tacvorian A Thesis Submitted to the Faculty of the WORCESTER POLYTECHNIC INSTITUTE In partial fulfillment of the requirements for the Degree of Master of Science In Biomedical Engineering November 6, 2012 APPROVED: Glenn Gaudette, Ph.D., Advisor, Biomedical Engineering Kristen Billiar, Ph.D., Biomedical Engineering Michael Kowaleski, DVM, DACVS, Cummings School of Veterinary Medicine

2 Acknowledgements Many thanks to my committee members Professor Kristen Billiar and Michael Kowaleski, DVM, DACVS, for their direction, support, and continued guidance throughout my graduate education. I would like to thank Jeremy Skorinko and Harry Hovagimian for their help throughout the testing process. I would also like to thank Julien Cabassu, DVM, Cara Blake, DVM, Randy Boudrieau, DVM, DACVS, and (again) Michael Kowaleski, DVM, DACVS, from the Cummings School of Veterinary Medicine at Tufts University for their clinical expertise and assistance which have allowed for the completion of this research. Without the continued support and encouragement from my mother, father, brother, family, and friends, I would have accomplished nothing. Finally, I would like to extend my extreme gratitude to Professor Glenn Gaudette for his invaluable guidance and countless hours of support throughout my graduate education and research. ii

3 Abstract The understanding of bone healing and principles of fracture fixation have improved greatly over the past fifty years. Plating systems are ideal for use in fracture fixation as they facilitate direct and indirect bone healing due to the stability they provide at the fracture site. Their main failure mode, however, is through fatigue from the consistent loading and unloading of the plated bone when healing. The goal of this study was to evaluate the mechanical properties of the most prominent veterinary plating systems representing a comminuted fracture when mated to a bone model. These assemblies were loaded to acute failure in four-point bending and cycled in torsion to mimic fatigue loading. Based on the analyzed test data we are able to make a number of conclusions. After performing four-point bending tests, the String Of Pearls (SOP) system sustained the highest bending mechanical properties with a bending stiffness of 80.4±12.5 N/mm, bending structural stiffness of 8.7±1.4 N-m 2, and bending strength of 11.6±1.7 N-mm. The Advanced Locking Plate System #10 (ALPS10) sustained the lowest bending mechanical properties with a bending stiffness of 40.0±1.9 N/mm, bending structural stiffness of 4.3 ± 0.2 N-m 2, and bending strength of 5.1±1.2 N-mm. Analysis of the cyclic fatigue data allow us to conclude that the Dynamic Compression Plate (DCP) system is able to maintain the highest absolute torque value across 15,000 torsion cycles and Fixin the lowest. This translates to 5.4±0.7 N-m and 3.5±0.4 N-m, respectively, when analyzed with Dixon-Mood equations and 5.4±2.5 N-m and 3.5±1.3 N-m, respectively, when analyzed with probability plots. In addition, the ALPS10 system is able to maintain the highest percentage of its failure torque and SOP the lowest. This translates to 76.4±16.3% and 43.6±5.3%, respectively, when analyzed with Dixon- Mood equations, and 72.9±28.6% and 44.2±22.1% when analyzed with probability plots. To aid in proper fracture healing, plating systems offering reduced or no contact with bone when applied in addition to screw holes across the entire plate length are preferred. The results of this evaluation are a start to better understanding plating system mechanics, which to develop further, will require further fatigue life testing in both loading conditions. iii

4 Table of Contents Acknowledgements... ii Abstract... iii 1. Introduction Background Bone Bone Anatomy Biology of Fracture Healing Unstable Fractures Inflammatory Phase Repair Phase Remodeling Phase Healing Under Restricted Motion Stable Fractures Contact Healing Gap Healing Bone Response to Mechanical Loads Stress Shielding Fracture Types Plating Systems Screws Self-Tapping Screws Standard Screws Locking Head Screws Plates DCP LC-DCP LCP SOP Fixin ALPS Fixing Plates to Bone Bone Loading Tension Compression Bending Torsion Plating System Load Distribution Bridging the Gap Goals Specific Aim Specific Aim Experimentation Bone Model Plates and Screws Assembly tools Aim #1 Experimentation Sample assembly iv

5 4.4.2 Initial Mechanical Testing Final Testing Aim #1 Results Aim #2 Experimentation Sample Assembly Initial Mechanical Testing Final Testing Results Results Four-point Bending Torsion Discussion Four-Point Bending Torsion Cinical Relevance Conclusions Bibliography Appendix A: Four-Point Bending Force-Displacement Graphs Appendix B: Four-Point Bending Calculated Data Appendix C: Four-Point Bending Minitab Analysis Appendix D: Cyclic Torsion Staircase Data per Plating System Appendix E: Cyclic Torsion Total Angular Rotation and Failure Modes Appendix F: Cyclic Torsion Probability Plots Appendix G: Dixon-Mood Calculations Appendix H: Plot of Runout Torque vs. Rotational Displacement v

6 Table of Figures Figure 1: Structure of an Osteon[5]... 3 Figure 2: Microscopic Anatomy of Compact Bone.[5]... 4 Figure 3: Gross anatomy of the long bone. (a) Long bone structure. (b) Cancellous bone structure. (c) Cortical bone structure.[5]... 5 Figure 4: Phase Timeline of Spontaneous Healing.[5]... 8 Figure 5: Cutting cones in stable fracture healing.[5] Figure 6: Microradiographs of plated femora. (a) SS plate (b) PTFCE plate.[25] Figure 7: Fracture Types: (a)transverse (b)oblique (c)spiral (d)comminuted (e)segmental Figure 8: Conventional Screw.[32] Figure 9: Locking Head Screw.[32] Figure 10: Dynamic Compression Plate Figure 11: Low Contact Dynamic Compression Plate Figure 12: Locking Compression Plate Figure 13: String of Pearls Plate Figure 14: Fixin Plate Figure 15: Advanced Locking Plate System # Figure 16: Advanced Locking Plate System # Figure 17: Eccentric Drill Guide (Left). Universal Drill guide (Right) Figure 18: Screw Securing Orders "A" and "B" Figure 19: Bone Loading. (a) Axial tension/compression. (b) Bending. (c) Torsion. (d) Shear.. 25 Figure 20: Axial and Bending Load Distribution for Conventional Plating Systems.[5] Figure 21: Axial and Bending Load Distribution for Locked Plating Systems.[5] Figure 22: Torsion Load Distribution for Conventional Plating Systems. Adapted from [47] Figure 23: Torsion Load Distribution for Locked Plating Systems. Adapted from [47] Figure 24: Bone Model Cross-Section Figure 25: Drilling Guide and Spacer Figure 26: Centering Jig Figure 27: Fixin Screw Deflection Figure 28: Properly Installed Four-Point Bending Test Sample Figure 29: Anvil Span Measurement Locations. Adapted from ASTM F382.[7] Figure 30: Sample Load-Displacement Curve Figure 31: Complete Torsion Bone Model Segment (a) Front View (b) Top View Figure 32: Initial Torsion Test Setup Figure 33: Positive Cycle Data for ALPS11 Cycled at 2Hz Figure 34: Assembled Bone Segment End Figure 35: Sample Plating System run-out and Failure Summary Figure 36: Sample Probability Plot Figure 37: Four-point bending plate deformation. From left to right: DCP, LCP, SS LC-DCP, Ti LC-DCP, ALPS11, ALPS10, Fixin, and SOP vi

7 Figure 38: Fixin Bone Model Fracture and Plate Deformation at Plate/Bone Model Interface Figure 39: Fixin Bone Model Failure Figure 40: ALPS10 Progressive Failure Figure A.1: ALPS10, Sample Figure A.2: ALPS10, Sample Figure A.3: ALPS10, Sample Figure A.4: ALPS10, Sample Figure A.5: ALPS11, Sample Figure A.6: ALPS11, Sample Figure A.7: ALPS11, Sample Figure A.8: ALPS11, Sample Figure A.9: DCP, Sample Figure A.10: DCP, Sample Figure A.11: DCP, Sample Figure A.12: DCP, Sample Figure A.13: Fixin, Sample Figure A.14: Fixin, Sample Figure A.15: Fixin, Sample Figure A.16: Fixin, Sample Figure A.17: SS LC-DCP, Sample Figure A.18: SS LC-DCP, Sample Figure A.19: SS LC-DCP, Sample Figure A.20: SS LC-DCP, Sample Figure A.21: Ti LC-DCP, Sample Figure A.22: Ti LC-DCP, Sample Figure A.23: Ti LC-DCP, Sample Figure A.24: Ti LC-DCP, Sample Figure A.25: LCP, Sample Figure A.26: LCP, Sample Figure A.27: LCP, Sample Figure A.28: LCP, Sample Figure A.29: SOP, Sample Figure A.30: SOP, Sample Figure A.31: SOP, Sample Figure A.32: SOP, Sample Figure D.1: ALPS10 Staircase Data Summary Figure D.2: ALPS11 Staircase Data Summary Figure D.3: DCP Staircase Data Summary Figure D.4: Fixin Staircase Data Summary Figure D.5: LCP Staircase Data Summary vii

8 Figure D.6: LC-DCP Staircase Data Summary Figure D.7: SOP Staircase Data Summary Figure F.1: ALPS10 Probability Plot, Percentage of Failure Torque Figure F.2: ALPS10 Probability Plot, Absolute Torque Value Figure F.3: ALPS11 Probability Plot, Percentage of Failure Torque Figure F.4: ALPS11 Probability Plot, Absolute Torque Value Figure F.5: Probability Plot, Percentage of Failure Torque Figure F.6: DCP Probability Plot, Absolute Torque Value Figure F.7: Fixin Probability Plot, Percentage of Failure Torque Figure F.8: Fixin Probability Plot, Absolute Torque Value Figure F.9: LCP Probability Plot, Percentage of Failure Torque Figure F.10: LCP Probability Plot, Absolute Torque Value Figure F.11: LC-DCP Probability Plot, Percentage of Failure Torque Figure F.12: LC-DCP Probability Plot, Absolute Torque Value Figure F.13: SOP Probability Plot, Percentage of Failure Torque Figure F.14: SOP Probability Plot, Absolute Torque Value viii

9 Table of Figures Table 1: Greyhound and Pit-bull Long Bone Mechanical Properties.[8]... 5 Table 2: UTS and Max Elongation of tissues throughout fracture healing.[5]... 7 Table 3: Plate and screw system features Table 4: Bone Segment Lengths Table 5: Sample Dixon-Mood Variable Calculation Table 6: Four-Point Bending Mechanical Property Summary Table 7: Four-Point Bending Tukey Analysis Summary Table 8: Plating System Torque Values Table 9: Plating System Failure Mode Occurrences Table 10: Mean Fatigue Strength Calculations Table 11: Dixon-Mood and Svensson-Loren Standard Deviation Comparison Table B.1: Proof Load Calculations Table B.2: Bending Stiffness Calculations Table B.3: Bending Structural Stiffness Calculations Table B.4: Bending Strength Calculations Table G.1: ALPS10 Dixon-Mood Calculations Table G.2: ALPS11 Dixon-Mood Calculations Table G.3: DCP Dixon-Mood Calculations Table G.4: Fixin Dixon-Mood Calculations Table G.5: LCP Dixon-Mood Calculations Table G.6: LC-DCP Dixon-Mood Calculations Table G.7: SOP Dixon-Mood Calculations ix

10 1. Introduction The understanding of bone healing and principles of fracture fixation has greatly developed over the past half century. Depending on the method of fixation, bone may heal in unstable or stable modes. An unstable fracture begins healing spontaneously, forming a protruding callus at the fracture site in the process. If a fracture is left to heal in this mode, it may take between six and nine years to fully complete the healing process. Fracture healing under stable fixation methods occurs without formation of callus, and reduces fracture healing times to about eighteen months. As controlling the movement and exercise of a canine with a fractured leg is difficult, stable fixation is preferred. Medical devices for human use must meet numerous requirements and regulations set by the Food and Drug Administration (FDA) to ensure their safety and effectiveness. While the FDA recommends that manufacturers of veterinary devices conduct tests to ensure their safety and effectiveness, there are no regulations governing their approval for use. Consequently, very few studies have been conducted calculating different mechanical properties of fracture fixation systems and assessing their similarities and differences. Studies researched varied in their testing methods as analyzed fixation devices were limited in addition to a restricted number of studies utilizing in vivo loading considerations. This limitation prevents the surgeon from determining the preferred fixation mode.[1-3] Forces must be applied to bone to facilitate healing. When fixators are used to stabilize bone fracture fragments, it is important for the stiffness of those fixators to be similar to that of bone. As stated by Wolff s law, bone undergoes many adaptations throughout its life to adapt to its mechanical environment. As such, if fixators absorb a majority of the load placed on the bone, stress shielding will effectively cause the fragment ends to resorb. This can lead to delayed union, nonunion, or lack of bone growth. Fixation plates are used for the stabilization of fractures in animals and humans. Using the canine femur as a model, there exist numerous principal plate systems, of which the most prominent were evaluated. Each of these systems contains a plate to span the fracture gap and corresponding screws to affix the plate to the bone. These systems vary in their plate dimensions and geometry, screw type, screw hole quantity, healing mode of plate application ultimately affecting the effectiveness, stiffness, and longevity of the system when placed under acute and cyclic loads. Using a synthetic bone model and set fracture gap, the plating systems were subjected to experiments to determine their bending strength under acute four-point bending loads and their fatigue strength under cyclic torsion loads, as excess loading in these modes are typically responsible for bone fracture. 1

11 2. Background Bones are comprised of numerous tissues, vessels, and chemical structures and serve as a rigid frame for the body s tissues while also protecting internal organs from impact forces.[4] More importantly, fracture or other damage should not occur from strain caused by repetitive everyday activities. However, depending on an activity s intensity level, duration, and repeated loading, microscopic damage may be seen.[5] While osteoclasts and osteoblasts work together to repair and maintain the structural integrity of bone, any damage occurring faster than the rate of repair will experience fracture. Bone fracture may heal through different methods depending on the damage level or fixation stability. Spontaneous healing occurs in fractures that are not fixated in a stable manner, while primary healing occurs in rigidly fixed fractures.[5] As this study focuses on femoral fracture fixator evaluation, the main fracture types associated with the diaphysis of the femur are transverse, oblique, and comminuted. Numerous fixation methods including intramedullary nails, pins, lag screws, and plating systems can be used to fixate these fractures, in addition to allowing them to heal spontaneously without fracture stabilization. 2.1 Bone The structure of bone can be broken down into multiple levels. While osseous tissue dominates the makeup of bone, numerous structures exist on microscopic and chemical levels. Epiphysis refers to the proximal and distal ends of bone while diaphysis refers to the main shaft. A majority of the epiphysis is comprised of cancellous bone while the majority of the diaphysis is comprised of cortical bone Bone Anatomy Mammalian bone, including that of humans and canines, is a composite material consisting of an organic matrix and inorganic hydroxyapatite. The wet weight of bone is derived 10-20% from water, 45-60% from hydroxyapatite (HA), and 30-35% from organic substances. The organic composition can be further broken down to 90-95% collagen, 1% glycosaminoglycans, and 5% other proteins.[6] There are four levels associated with the structure of bone. Fundamentally, HA crystals are ingrained between the ends of adjacent collagen fibrils. When separate, HA and collagen do not possess high mechanical properties, but their combined form yields a composite with excellent mechanical properties[6]. Generally, bone is more ductile than HA and is able absorb more energy before failure and bear higher loads as it is more rigid than collagen. On the second level, lamellae form from the combined collagen-ha fibrils and have specific orientations which define the strength limits in their primary loading direction. This arrangement of lamellae is seen on the third level. A tubular Haversian osteon is one functional unit of bone, produced from the circular and concentric lamellae structure and possesses its maximum strength along the long axis.[5] This is seen in Figure 1. 2

12 Figure 1: Structure of an Osteon[5] The macroscopic structure of bone is observed on the fourth level of bone organization. At this level, the main factors determining bone s strength are its density and trabecular orientation. There are two main bone structure types: cortical and cancellous. Cortical, or compact, bone is found mainly in the diaphysis of long bones and the exterior shell of other bone types. Compact bone consists of Haversian osteon systems and interstitial bone regions. Osteons are typically oriented in the longitudinal direction of long bones and are typically 200µm in diameter and 10-20mm long. They are further composed of concentric lamellae 3-9 µm thick.[6] Haversian canals run through the center of osteons and allow blood vessels to deliver nutrients to osteocytes (bone cells). Biomechanically, cortical bone can be characterized as being semibrittle, viscoelastic, and its strength is orientation dependent. Figure 2 shows the microscopic anatomy of compact bone. 3

13 Figure 2: Microscopic Anatomy of Compact Bone.[5] Cancellous, or trabecular, bone contains a highly porous structure and is located on the epiphysis of the bone. Pores are interconnected and filled with bone marrow, making up 75-95% of the bone volume. Cancellous bone is made up of a matrix of small struts and plates called trabeculae that are between µm thick and are spaced µm apart.[6] This porous structure of cancellous bone has distinctly different mechanical properties from cortical bone. The structural and apparent densities, strength, and moduli of cortical bone are all considerably greater than cancellous bone. Stiffness and strength are the two core mechanical properties of bone. Stiffness is expressed by the elastic modulus and is calculated by the stress required to elastically deform bone. Deformation, or change in shape, occurs in structural materials as they are loaded. If this change in shape reverses with the removal of the load, the material is said to have undergone elastic deformation. If this change in shape is permanent, plastic deformation has occurred. Bone strength is defined as the stress required to cause plastic deformation or fracture.[7] Various mechanical tests can be performed to determine the stiffness and strength of bone. These measurements can be recorded and calculated using a load-displacement curve. Table 1: quantifies the mechanical properties of greyhound and pit-bull long bones as previously examined by Kemp et al.[8] Figure 3 displays the gross structure of a long bone. 4

14 Table 1: Greyhound and Pit-bull Long Bone Mechanical Properties.[8] Greyhound (28.52±1.98 kg) Pit-bull (23.61±3.73 kg) Elastic Modulus (GPa) Humerus 7.70± ±0.34 Radius 15.07± ±0.68 Femur 11.22± ±0.62 Tibia 14.05± ±0.44 Yield Stress (MPa) Humerus ± ±28.75 Radius ± ±10.86 Femur ± ±10.14 Tibia ± ±2.90 Figure 3: Gross anatomy of the long bone. (a) Long bone structure. (b) Cancellous bone structure. (c) Cortical bone structure.[5] 5

15 2.2 Biology of Fracture Healing When bone experiences a load exceeding its ultimate tensile strength, fracture occurs. Fracture is defined as a breach in continuity of bone, either on a macroscopic or microscopic level. Fractures may heal via indirect or direct healing, depending on their relative stability Unstable Fractures Unstable fractures heal via indirect healing, which is characterized by the formation of callus as a process intermediate before modeling into hard bone.[5] This is also the mode of healing for nonoperative fracture treatment, unstable internal and external fixation, along with the plate osteosynthesis of highly comminuted fractures.[9, 10] The amount of callus produced is inversely proportional to the stability of the fraction, as a less stable fracture results in increased callus formation, and vice versa.[11] This is dictated by interfragmentary strain. Interfragmentary strain is the deformation which occurs at a fracture s bone fragment interface. This is a major influence in the progression of fracture healing and is calculated by dividing the displacement of the fracture gap by the initial gap width. Bone and callus formation cannot occur with an interfragmentary strain greater than two percent. To overcome high strain, muscles surrounding the bone first contract to begin resorption of the fragment ends. At strains between two and ten percent, a fibrocartilage forms and between ten and one-hundred percent a granulation tissue forms.[12] Once the tissues surrounding the fragment provide an interfragmentary strain below two-percent, callus formation may begin.[5] The timeline of unstable fracture healing is divided into three overlapping phases: inflammatory, repair, and remodeling Inflammatory Phase Once the integrity of bone and its surrounding tissues are disrupted, the inflammatory phase begins and continues until the initiation of cartilage or bone formation. This typically lasts three to four days depending on the level of damage and magnitude of force causing the bone disruption.[5] Within hours after bone disruption, an extraosseous blood supply emerges from the surrounding tissues to begin the revascularization of the hypoxic fracture site. This forms a fibrin rich clot and initiates spontaneous healing. Growing evidence suggests that hematoma fosters the repair phase by releasing growth factors to stimulate angiogenesis and bone formation.[5] Vasoactive substances are released by mast cells and are believed to contribute to new vessel formation.[13] Macrophages also play a role in fracture repair as they release fibroblast growth factor (FGF), initiating fibroplasia both in soft tissues and in bone. As vasculature is reconstructed, internal, or medullary, blood flow resumes and the extraosseous blood supply diminishes. Mesenchymal stem cells (MSCs) from the cambium layer of the periosteum, endosteum, bone marrow, and adjacent soft tissues proliferate during this phase.[9] Unless infection, excessive motion, or extensive necrosis of the soft tissue is noted, the hematoma will resorb by the end of the first week after bone disruption. The end of this phase is marked by a decreased observation in pain or swelling. 6

16 Repair Phase During the repair phase, hematoma is transformed into granulation tissue with assistance from capillary ingrowth, mononuclear cells, and fibroblasts. While interfragmentary strain remains high in this phase, the sustained formation of granulation tissue is explained by its ability to stretch to twice its length. The first observation of mechanical strength in the disrupted bone is observed during this phase as the formed granulation tissue has a tensile strength of 0.1 N/m 2. Granulation tissue helps reduce interfragmentary strain at the fracture site and in turn initiates cartilage formation.[5, 11] The tissue matures predominantly into Type I collagen fibers that have an ultimate tensile strength of 1-60 N/m 2 and can resist elongation to a maximum of 17%.[5] As the fibers mature, they organize into a diagonal pattern to allow for optimized elongation ability. Many elements including low oxygen tension, poor vascularity, growth factors, and interfragmentary strain influence the ability of the collagen fibers to develop into a cartilaginous callus. Proliferated MSCs present from the inflammatory phase are orchestrated by TGF- and morphogenic proteins (BMPs) to differentiate into chondrocytes. This differentiation is essential to the maturation of collagenous fibers at the fracture gap into internal and external cartilaginous soft callus matrices, which are facilitated by angiogenesis and an intact periosteum. In well-vascularized unstable fractures, a bulging external callus is also found, increasing the injured bone s resistance to bending. An increase in proteoglycan concentrations in the fibrocartilage is observed, contributing to increased stiffness at the interfragmentary gap. This callus resists compression, but has a similar ultimate tensile strength and elongation before rupture as fibrous tissue (4-19 N/m 2, %). Once interfragmentary strain reduces to below ten percent, this cartilaginous matrix may mineralize, maturing into hard callus by endochondral ossification.[9, 11] In this process, chondrocytes calcify and degenerate, facilitate angiogenesis, and allow osteoblasts to lay down woven bone on the collagen framework left by the chondrocytes. The length of time necessary to achieve union depends on the fracture configuration and location, the status of adjacent soft tissues, in addition to the patient s statistics. While bone union is achieved at the end of the repair phase, its existing structure at the fracture site does not resemble that of the original bone. At the end of the repair phase, however, enough strength and rigidity has been regained to allow a canine to begin low impact exercise. Table 2 describes the ultimate tensile strength and maximum elongation of tissues formed throughout the fracture healing process. Table 2: UTS and Max Elongation of tissues throughout fracture healing.[5] UTS (N/m 2 ) Max Elongation Granulation Tissue % Collagen Fibers % Early soft callus % Bone 130 2% 7

17 Remodeling Phase In the final phase of fracture repair, hard callus undergoes a morphological adaptation to regain the optimal function and strength of intact bone. In humans, 6-9 years can pass until the process completes, and represents 70% of the total healing time. During this phase, the woven hard callus remodels into the required amount of lamellar bone and any excess callus is removed, thereby restoring the medullary canal and bone shape.[11] Bone is arranged in areas experiencing excess stress and removed from areas where there is too little, according to the accepted theory known as Wolff s Law which states that bone structure constantly adapts to the mechanical loads to which it is subject. Figure 4 shows the phases and time distribution required to spontaneously heal bone. As previously described, the progression of soft to hard callus in spontaneous fracture healing is dependent on the fracture site being supplied with sufficient blood in addition to consistent increase in stability. Without these two factors, fibrous tissue will not advance to hard callus and will lead to an atrophic nonunion. This is typically remedied with the addition of bone grafts or removing the layer of both on the two apposed fracture ends to restart the healing process. If proper vasculature exists without interfragmentary motion control, the fracture will progress into a cartilaginous callus unable to stabilize the fragments and will further progress into hypertrophic nonunion. This is typically remedied with the addition of rigid fixation. When canine bone undergoes spontaneous bone repair, malunion is not uncommon.[5] Figure 4: Phase Timeline of Spontaneous Healing.[5] Healing Under Restricted Motion Fractures controlled under restricted motion heal in a process intermediate to spontaneous healing of uncontrolled fractures and healing after absolute stabilization. To limit motion at the 8

18 fracture gap and minimize the likelihood of malalignment, pins and nails are often implanted. Healing of fractures under restricted motion begins with some resorption of the fragment ends. Primitive implantation methods advised reaming the medullary canal prior to implanting nails, thereby removing bone marrow and disrupting endosteal blood flow. This allowed for the largest possible nail diameter to be implanted in the medullary canal. The initial stability of these implants is attributed to the contact between the nail and the bone s inner cortex, in addition to the screws used to provide rotational stability. The process of reaming prior to implantation, however, decreases the blood supply available at the fracture site by 70%.[14] Since research has established that an appropriate blood supply is required for fracture healing, nailing methods and implants have been modernized to minimize disruption. This includes making optional the reaming of the medullary canal in addition to changes to implant geometry, which together have demonstrated only a 30% reduction in blood supply at the fracture site. Unreaming nails, however, do not offer the initial fixation stability of reaming nails. Some research has shown that the increased blood supply associated with unreamed implants may not correspond to improved healing times.[15] While the restricted motion from IM nailing demonstrates a significant improvement over spontaneous healing, the ossification associated with healing under restricted motion is only about 10% of that associated with plate or external fixator stabilization.[15] Advocates of unreamed nailing believe reaming of the intramedullary canal is detrimental as the endosteal blood flow will be disrupted and may potentially further damage the bone.[15, 16] Though studies have debated the healing success of reamed versus undreamed intramedullary nails, stable fixation remains the most effective form of fracture management as it facilitates direct healing. External fixators may also be used to restrict motion and are applied using closed reduction techniques while further minimizing vasculature disruption and maintaining stability. The amount of callus formed is minimal but can vary greatly depending on the configuration of the fracture and the rigidity of the fixator frame applied.[5] Variation can occur if the implant is not placed on the tension side of the bone, fracture reduction is not perfect, or if the implant lacks rigidity. These factors are most relevant in fixation of comminuted fractures since fragment ends are more difficult to align properly and mechanical stability greatly influences the course of healing. While closed reduction external fracture fixators offer decreased callus formation, their structure may not provide adequate stability due to the moment created by implant being offset from the body. More importantly, a higher probability of infection exists as the implant must pass through the patient s dermis in multiple locations. Overall, while healing by restricted motion may pose benefits to the healing process over spontaneous healing, increased stability and lack of callus formation resulting from stable fracture fixation is optimal Stable Fractures In 1949, Danis reported that a callus is not formed during bone healing when two bone fragments are apposed under a rigid plate and axially compressed.[5] Application of rigid, nongliding implants, such as compression plates and lag screws results in a stable bone fragment 9

19 interlocking connection. It was later confirmed by Schenk and Willenegger that healing under these conditions was the result of direct osteonal proliferation.[5, 17] This repair mode, termed primary or direct healing, refers to direct filling of the fracture site with bone, without the formation of periosteal or endosteal callus. Rigid fixation and precise reduction suppress the biological signals found through indirect healing methods which attract callus promoting osteoprogenitor cells to the fracture site. Interfragmentary compression induced by a plating system creates differences in the biomechanical microenvironment within the fracture site and influences the progression of bone development.[5] Contact and compression areas are bounded regions separated by areas where fragments ends are separated by small gaps. A compression plate applied across a fracture site will create two different healing zones. The cortex directly under the plate will experience contact and compression characteristics triggered by the plating system. The far cortex will be exposed to forces in tension and will be subject to gap healing. Both contact and gap healing are mechanisms classified under direct fracture healing. Utilization of plating systems to facilitate primary healing remains the method of choice when fracture healing is required.[5] Contact Healing Contact healing in stable fractures is observed when the defect between bone ends is smaller than 0.01mm and the interfragmentary strain is less than 2%.[3, 5, 9, 12] Lamellar bone is directly formed as a result of primary osteonal reconstruction and is oriented in the normal axial direction. Cutting cones are cells which form at the ends of the osteons closest to the fracture site. These are seen in Figure 5, adapted from Rüedi et al.[5] Osteoclasts line the head of the cutting cones while osteoblasts line the tail. This enables bony union and Haversian remodeling to occur simultaneously.[5] Osteoclasts advance across the fracture site and create longitudinally oriented cavities in which the osteoblastic ends deposit osteoids. Cutting cone navigation has been reported across canine radial osteotomies under rigid fixation as early as three weeks after surgery. Cutting cones travel across fragments at a rate of µm/day and become the spot welds which unite the fragment ends without the production of callus.[17] The bone formed during remodeling will be visible on radiographs until bone density at the fracture site reaches that of intact cortical bone Gap Healing Direct healing at gaps between 800µm to 1mm occurs in a similar process to contact healing, though bony union and Haversian remodeling remain separate, sequential steps.[9] Healing starts with osteoblasts depositing layers of lamellar bone on both fracture surfaces, perpendicular to the long axis, until the ends unite.[9] This area, however, remains weak due to this bone orientation. Haversian remodeling initiates between three and eight weeks after surgery when osteoclasts form on both fracture ends and create longitudinally oriented resorption cavities. Longitudinally oriented lamellar bone is deposited into these cavities over time by the osteoblast tail of the cutting cone so anatomical and mechanical integrity of the cortex may be reestablished.[5] 10

20 Figure 5: Cutting cones in stable fracture healing.[5] (a) Gap healing. (b) Contact healing. (c) Osteoclast head. (d) Osteoblast tail Bone Response to Mechanical Loads Contact and gap healing are facilitated by load transferred to the bone fragments. In 1892, Julius Wolff reported his research to explain how the orientation of trabeculae in the femur is established during development.[18] His findings matured into a general means to understanding the gross shapes and adaptations of bone. Often referred to as Wolff s Law, bones constantly grow and remodel to adapt to their mechanical environment.[4, 18-21] The primary function of bone is to remain stiff and resist deformation from internal muscular and external forces.[18] To maintain its stiffness, bone strength can be increased with added bone mass, a changed geometry to redistribute the acting forces, or altering its microstructure through Haversian remodeling. Intrinsic factors including osteon density, mineral density, porosity, collagen fiber orientation, and histologic structure affect the strength of bone. Loading mode, duration, and rate of strain are extrinsic factors corresponding changes in bone strength. The combination of these intrinsic and extrinsic factors affects bone s mechanical properties and adaptations in response to loading. Bone is a viscoelastic material which responds to mechanical loads differently depending on their magnitude. Wolff s hypothesis that osteocytes act as strain receptors and transducers has received the most attention in research.[18] When bone experiences a strain of sufficient magnitude to elicit a response, one of four outcomes may result. No osteogenic response may occur if the strain transduces a signal that is below a certain threshold or the receptor is inhibited as a result of aging. A sufficient signal will result in osteoblasts recruited in the periosteum or endosteum to initiate remodeling or osteoclasts recruited along the bone surface to initiate resorption. Finally, osteoclasts and osteoblasts can be activated sequentially to initiate Haversian remodeling, as described in Section [18] 11

21 Though the methods of sensing mechanical loads by bone cells are not well understood, there are many indications that strain rate and magnitude are important stimuli effecting bone response.[18, 21, 22] One of the hypotheses regarding the mechanism of mechanotransduction which has received the most attention suggests that osteocytes are mechanosensors of shear stresses.[23] Osteocytes have radiating canaliculi which communicate with other osteocytes through transmitter proteins at gap junctions. Together, these osteocytes form a connected cellular network surrounding the periosteal and endosteal membranes. These cells further connect to osteoblasts lining the periosteum and endosteum which connect to preosteoblasts in the membrane. Together, this network effectively creates a nervous system in the bone which controls nutrient flow and initiates bone remodeling when activated.[18] Stress Shielding As described by Wolff s law, mechanical loading of bone stimulates the initiation of Haversian remodeling. One issue observed with the introduction of rigid fixation systems is bone refracture after implant removal. Researchers have attributed this to bone atrophy as a result of the fixation system bearing a majority of the load, and not the bone.[24] This phenomenon has been termed stress shielding. Stress shielding is a common occurrence with rigid fixation systems and results in a loss of bone mass around the area where a plating system is applied.[12] The effects of this are apparent when a system is comprised of two or more components, as these components typically have different moduli of elasticity.[25, 26] An applied bone plating system, for example, creates a system comprised of the fractured bone, fixation plate, and plate screws as the components, which may all have different moduli of elasticity. When a load is placed on this system, the stiffer component bears more of the load, thus shielding the other components.[25] Research has suggested that fixation of fractured long bones with a plating system leads to osteoporosis and the possibility of fatigue fracture after its removal.[25] Immature bone formation and thinning of the cortical wall have been found at fracture sites shielded from loads by an apposed rigid plate.[25] As a system s material and geometry determine its stiffness, minimizing the effects of stress shielding will require a system s stiffness to be near that of bone. Preliminary clinical research, however, has shown promise in the use of internal fixators for stable fracture repair.[27] A study conducted by Tonino et al. in the early days of rigid fixation compared the effects of stress shielding in canine femurs by fixing six canines of similar weight with two different implant systems: the right femur with one comprised of stainless steel (SS) and the left femur with one of polytrifluormonochloroethylene (PTFCE). After harvesting and evaluating the femora, the femora fixed with the SS implant had significantly lower bone mineral mass per centimeter.[25] This occurred due to both bone resorption and only partial mineralization of the newly formed bone. Microradiographs of the cross-sectional area of bone healed with the SS system had larger resorption cavities, thinner lamellae and less woven bone formation than that of the PTFCE plate. This can be seen in Figure 6, adapted from Tonino et al. 12

22 (a) (b) Figure 6: Microradiographs of plated femora. (a) SS plate (b) PTFCE plate.[25] Mechanical testing confirmed these observations impacted the bone s strength as the femora healed with SS plates required a 30% lower force to fracture, 22% lower bending strength, and 20% lower modulus of elasticity than those healed with the PTFCE plate. The observed histological and mechanical differences are entirely due to the plate material as both systems had the same geometry and application area. It can therefore be concluded that the material composition of the SS plating system played a role in the stress shielding effects observed by the femora. Similar results have been reported by Diehl and Mittelmeier who observed a loss of function in tibias healed with stainless steel plating systems. They found the load required to fracture the bone to be one-third that of intact bone.[25] The effects of stress shielding are further amplified with systems disrupting bone s surrounding vasculature as this prevents bone growth beneath the plate. Recent plating system developments have attempted to address the issues surrounding stress shielding by changing plate material and geometry Fracture Types Break in the continuity of bone is classified as a fracture. When caring for the diaphysis of long bones, two fracture types may be observed, undisplaced or displaced.[28] In undisplaced fractures, bony fragments are still in their anatomical position, have cracks present, and do not require any reduction.[28] Displaced fractures may be further classified into five categories: transverse, oblique, spiral, comminuted, and segmental. Transverse fractures are found perpendicular to the long axis of bone and may occur due to numerous factors. Failure may occur under tensile or bending loads, a direct strike to the bone, or an indirectly delivered force as may be seen from a fall from significant height. Trauma to the bone may result in the fracture becoming more comminuted with progressively greater force.[28, 29] Oblique fractures are characterized by an oblique line found at from the long axis and are typically the result of combined bending and torsional forces.[28, 29] Bone failure as a result of torsional forces is a test of its mechanical properties in shear and tension. The torque moment creates a state of pure shear between parallel transverse planes and tension and compression forces at all angles in between. These forces both come to a maximum at 45 to the long axis of the bone, which when great enough to produce a fracture, result in a spiral shaped line.[29, 30] This type of fracture may be observed, for example, in cases where a canine paw becomes lodged while running, 13

23 indirectly distributing a torsional load to its long bones. A comminuted fracture is one where more than two bone fragments exist, typically including small wedges and small fragments which are nonreducible.[29, 31] Segmental fractures are a type of comminuted fracture where the fragments are whole and large enough to be anatomically reconstructed.[5] Figure 7 depicts these fracture types. Figure 7: Fracture Types: (a) Transverse (b) Oblique (c) Spiral (d) Comminuted (e) Segmental. 2.3 Plating Systems Screws and plates are used to achieve bony union of two or more fragment ends of a fracture, whether surgery is performed with Open Reduction Internal Fixation (ORIF) or Minimally Invasive Plate Osteosynthesis (MIPO) techniques. Depending on the type of fracture, different plates may be used to facilitate the proper healing function. Plates may accommodate conventional or locking screws. Multiple screw types may be used to attach a plate to bone. These include cortical and cancellous screws with self-tapping or standard threads, and may have locking heads. There are two plate types which can be applied: conventional and locking Screws When choosing screws for plate application, veterinarians are recommended to utilize a screw diameter should not exceed 40% of the fractured bone diameter to prevent a decrease in bone strength.[5] Conventional or locking head screws may be chosen for implantation depending on the plate being applied. A standard screw is depicted in Figure 8 and a locking head screw in Figure 9. Conventional screws are adapted to accommodate both cancellous and cortical bone. Cancellous screws have a relatively thin core with wide and deep threads, while cortical screws have a relatively thicker inner core with shallower threads. The increased ratio of the outer diameter to inner core of the cancellous screws allow for a significantly greater holding power in the trabecular bone of the metaphyses and epiphyses. Cortical screws are typically used in bone diaphyses. They can be fully threaded when used to fix plates to bone or partially 14

24 threaded when used as a lag screw. Lag screws are used when interfragmentary compression between bone fragments is required.[5] Self-Tapping Screws Self-tapping screws are designed to be screwed into bone after a pilot hole has been drilled. Self-tapping screws cut a thread into bone as they are being fastened. While self-tapping screws may be removed and reinserted, they are best used in applications where applied only once since an inadvertent misalignment after removal will destroy the previously cut thread and may cause premature failure of the plate Standard Screws Standard screws are used in conventional plates and when a need to replace or reposition the screw along the healing process is anticipated. A pilot hole is first drilled into the bone and then threads are cut into the hole using a tap corresponding to the threads on the screw being used. Standard screws may be removed and reinserted with ease and without fear of inadvertent thread damage Locking Head Screws Locking head screws may have standard or self-tapping threads on their core in addition to having threads surrounding the head of the screw. The screw head locks into the plate hole threaded to accommodate them. The threads on the head incorporate a different pitch and diameter than those on the core and thus provide a greater resistance to pullout and the ability remove any compressive forces between the plate and the bone. This is essential in preventing the disruption of vasculature around the affected site. Furthermore, as plates may sometimes be contoured and angled to better fit their application, these screws guarantee the plate s location is undisturbed. Figure 8: Conventional Screw.[32] Figure 9: Locking Head Screw.[32] 15

25 2.3.2 Plates Plates are designed to facilitate one or more of the following functions in fracture fixation: compression, neutralization, bridging, or buttress. Compression plating generates axial forces by use of a tensioning device or eccentrically loading screws. This mode is typically used in simple transverse fractures and those with low obliquity. If a diaphyseal fracture is fixed with a plate and screws to not produce any compressive axial forces, the system functions in a neutralization mode.[33] In addition, as lag screws may be used to fix fragments in comminuted fractures, a plate applied in neutralization mode protects the interfragmentary compression of the fragments from any rotational, bending, or shears forces when loaded.[5] A plate functioning as a buttress is applied to metaphyseal fractures to prevent the collapse of fragments when the articular surface is exposed to compressive forces.[5] A plate functioning as a bridge is applied when indirect reduction of bone is required, as in comminuted fractures. It functions as a splint to maintain correct length of the bone and normal joint alignment when fixing the fragment ends as it prevents axial deformity as a result of shear or bending forces.[5] Since the plate gets subjected to full weight-bearing forces, it is important that the soft tissue surrounding the fragments maintain their vascular supply as the success of indirect reduction is dependent on the formation of the bridging callus.[5] In all plates depicted below, the top panel shows the top surface of the plate, and the bottom panel shower the underside surface DCP The dynamic compression plate (DCP) was first introduced in 1969 and featured unique hole geometry which allowed for axial compression by eccentric screw insertion and is available in stainless steel.[5, 34] The success of bone healing utilizing this plate is dependent on the friction between the plate and the bone to generate a rigid internal fixation. This plate may facilitate healing modes by compression, neutralization, bridging, or buttress. When the screw is inserted into the plate and tightened, the bone fragment moves relative to the plate and consequently compresses the fracture ends axially. The shape of the holes allows for a 25 inclination of the screws longitudinally, and 7 transversely.[5] DCP s plate holes are symmetric and evenly distributed, allowing for versatile and eccentric placement of screws. This allows compression at any part of the plate and is advantageous in segmental fractures. Different drilling guides are available for different plate sizes and for the facilitation of eccentric and neutral screw loading. This plate can be seen in Figure 7. Recent studies have shown that implementing the DCP plating system may be detrimental to bone healing. Use of this system is now associated with a surgical technique that causes the plate to disrupt the blood supply underneath the plate, thus leading to delayed healing, nonunion, or an increased chance of bacterial infection.[35] Compromising the periosteal blood supply is a key disadvantage of using this system, and a factor which should be taken into consideration by the surgeon before deciding on an implant system. 16

26 Figure 10: Dynamic Compression Plate LC-DCP The low contact dynamic compression plate (LC-DCP) seen in Figure 11 is an advancement of the DCP system and is available both in stainless steel and titanium.[5] The LC- DCP was developed to alleviate some of the disruption to the periosteum seen with the DCP system. The geometry of the LC-DCP system accomplishes this by allowing 50% less contact with the bone that would otherwise be achieved with the DCP plate.[35] This geometry improves cortical perfusion and integrity of vasculature beneath the plate. The scalloped configuration of the plate s underside facilitates this and more evenly distributes load across the bone, allows contouring of the plate easier, and minimizes the possibility of damaging the screw holes when being contoured.[5, 35] When used in a buttress or bridging configurations, the plate geometry facilitates a distribution of load results in a minor elastic deformation of the plate without stress concentration at any of the screw holes, as would be present in DCP.[5] While this is an improvement from the DCP plating system, LC-DCP still requires compressive axial forces along the bone fragments and periosteum to facilitate proper healing. LC-DCP has similar characteristics to DCP. LC-DCP s plate holes are symmetric and evenly distributed, allowing for versatile and eccentric placement of screws. This allows compression at any part of the plate and is advantageous in segmental fractures. Screws may be placed up off-axis up to 7 transversely and 40 longitudinally, allowing more screw angulation freedom than DCP. The screws may be placed similarly to the DCP to facilitate compression, neutral, bridging, and buttress healing modes.[5] Similar to the DCP, various drilling guides are available for various screw loading techniques. As both DCP and LC-DCP plates both have the potential to disrupt vascularity, their efficacy may vary between patients. Figure 11: Low Contact Dynamic Compression Plate LCP The locking compression plate (LCP) is a unique implant as it incorporates the vascularity-preserving underside geometric advantages of the LC-DCP system while eliminating the need for compressive forces to be applied to the bone. This is achieved when using locking 17

27 head screws, though LCP is a combination hole plate which accommodates standard screws as well.[36] One-half of each hole is designed to accommodate the standard DCP and LC-DCP screws for fragment compression while the other half accommodates the locking head screws, advantageous for angular stability and removal of compressive forces from the bone fragment surfaces. As locking head screws have a larger core diameter, their use increases bending and shear strengths while displacing the load across a larger area across the bone. Use of locking head screws reduces the priority of perfectly contouring the plate due to the angular stability produced.[36] This plate is illustrated in Figure 12. In vitro biomechanical testing was conducted in a study by Aguila et al. comparing the LC-DCP and LCP plates when fixed to 14 pairs of femora with a 20mm osteotomy gap. No significant difference in structural stiffness of both plates was found in four-point bending. The LC-DCP system was found to be significantly stiffer when tested in cyclic torsion.[36] Figure 12: Locking Compression Plate SOP The string of pearls (SOP) plating system, illustrated in Figure 13, is a newer, stainless steel locking plate system designed both for veterinary and human use. Though it is a locking plate, it is secured using standard screws. Holes in the spherical pearls of the plate have threads which correspond to those on the core of the screws. As the screw is secured to the plate, it threads into the pearls and thus allows the screw to be properly torqued while removing any compressive force acting on the bone. The SOP system is similar to that of the LCP as they both locking plates which serve to minimize damage to the periosteal blood supply and minimize the need for plate contouring to the bone. The SOP plate is comprised of cylindrical internodes connecting the spheres, which have a greater moment of inertia over the DCP, LC-DCP, and LCP plating systems due to their geometry. The internodes are 5mm in diameter while the spheres are 8mm in diameter for all plate lengths. These cylindrical components allow the plate to have up to six degrees of rotational freedom when contouring is necessary.[37] The design of these components also prevents potential deformation of the screw holes when being contoured, a drawback to the flat locking plate systems. In a four-point bending study conducted by Ness, the SOP plating system attained a higher bending stiffness, bending structural stiffness, and bending strength than the DCP system.[37] Further testing was conducted comparing bent, twisted, and contoured SOP plates to the untouched DCP plate. The results of these tests demonstrated that the twisted and contoured SOP plate maintained higher strength and stiffness properties. No significant difference was 18

28 found between the mechanical properties of the bent SOP plate and the untouched DCP plate.[37] Another study conducted by Ness evaluated the outcome of humeral fracture repair in canines with a mean weight of 22.8 kg where two SOP systems were applied to the fracture. Postoperative analysis of the thirteen canine humeri demonstrated satisfactory function of the repaired limb in 12 of the 13 canines.[38] Additional surgery due to complications was recorded in four canines, three of which demonstrated satisfactory function after healing. Refracture was only evident in one canine. No screw loosening, backing out, or breakage was observed in the 115 SOP screws used for fracture fixation.[38] While the functional outcome following surgery was excellent in most cases, no bone density analysis was performed analyzing the healed bone. Figure 13: String of Pearls Plate Fixin Fixin is a novel locking plate system which incorporates a bushing insert between the screw-plate interfaces. It facilitates locking by creating a friction fit between the conical bushing and screw head as the bushing threads screw into the plate and the screw threads into the fractured bone. The bushing s titanium make-up allows for easier removal of the implant as the any concern of removal complications resulting from cold welding, cross threading, or damage to the hexagonal screw recess previously reported with other locking plate systems.[39] This combination of features allows the Fixin plating system to be angularly stable, simple to apply, and easy to remove when necessary. The Fixin plate is made of stainless steel and has threaded holes for the titanium bushing inserts. The screws used in this plating system are typically stainless steel, self-tapping, and are used in a locking mode. They have a larger core diameter to increase bending and shear strength while also improving load distribution along the bone. The head of the screw incorporates a conical surface matching that of the titanium insert, allowing stability through friction, microwelding, and elastic deformation.[39] While no previous studies have been found evaluating the stability and stiffness of this system, typical patients are canines and felines weighing up to 10kg. This is illustrated in Figure 14. Figure 14: Fixin Plate 19

29 ALPS The Advanced Locking Plate System (ALPS) is a novel system incorporating a uniform cross-sectional moment of inertia along the entire length of the plate due to its geometry where the screw hole sections are wider than those connecting them.[40] The profile of this plate also allows for improved periosteal blood flow in comparison to standard plates as there is minimal contact with the bone.[40] The scalloped geometry and titanium makeup allow for increased resistance to infection and deceased healing time as contact with the bone is decreased.[40] The screw holes on the plate allow for standard screws to be placed in various angulations, or locking head screws in fixed angulation. The locking mechanism of the hole functions by engaging the threads on the screw shaft. As the screw reaches its last few threads by the screw head, the thread diameter is reduced and the plate is pulled toward the bone when the locking mechanism is fully engaged.[40] The conical interface between the screw heads and plate hole also allows both screw types to be held in a stable position. The manufacturer of this plating system, Kyon Pharma Inc., claims the strength of their ALPS10 plating system is higher than that of the 3.5 DCP system.[41] In a study case report written by Inauen et al., an ALPS5 plate was implanted into a two-year old feline to repair the left hind limb lameness caused by a separation in an arthrodial joint of its foot. Five days postsurgery, the owner reported no apparent lameness, and unrestricted activity was allowed.[40] The surgeon had initially advised two weeks of cage rest followed by three weeks of restricted activity. Clinical examination six weeks after surgery proved the feline was without lameness and its left tarsus was stable and not swollen.[40] Furthermore, radiographs depicted uneventful healing. As a result of this study, the ALPS system is shown to be effective in conditions where standard compression plates may be applied for healing.[40] ALPS10 and ALPS11 are depicted in Figure 15 and Figure 16. Figure 15: Advanced Locking Plate System #10 Figure 16: Advanced Locking Plate System # Fixing Plates to Bone In summary, conventional plates require compressive forces to be applied to the bone fragments, require contouring to the bone, and may disrupt the bone s surrounding vasculature 20

30 depending on the geometry of the plate s underside. Locking plates do not require compressive forces to be applied to the bone and also do not require exact contouring to the bone to facilitate adequate healing and as a result of no compressive forces being applied to the bone, the bone s surrounding vasculature is preserved. However, due to their geometry and the method of application, locking screws cannot be applied eccentrically in these plates to provide compression at the fracture site. When plating systems are required to heal bone fracture, Open Reduction Internal Fixation (ORIF) or Minimally Invasive Plate Osteosynthesis (MIPO) techniques may be employed to secure plates to bone. To allow proper alignment, ORIF utilizes direct reduction techniques as the fracture site and surrounding tissues are exposed. This is typically achievable with transverse, oblique, segmental, and comminuted fractures with one to three reducible fragments.[5, 28] This treatment allows fragments to be perfectly aligned, but comes at the cost of increase healing time due to the disruption of surrounding tissues. Plates fixed to these fracture types will act in compression or neutralization modes. MIPO utilizes indirect reduction techniques to align fracture fragments without exposing the fracture site or disrupting the periosteum, thus maximizing the healing potential of the fracture site.[42] Reduction instruments are inserted through skin incisions made near the fracture to align the fragments.[42] After proper alignment, a plate is inserted through the incisions. Incisions are then made over the area of the plate where screws are inserted. This is typically performed in undisplaced or highly comminuted fractures.[42] Plates fixed to these fracture types will act in bridging or buttress. Fractures reduced by direct reduction will be healed via direct healing.[5, 9] As anatomical reconstruction is not possible in fractures reduced by indirect reduction techniques, these fractures will heal via indirect healing.[5, 9] The correct plating system to stabilize a fracture is dependent on the required fixation mode. Plate size is typically dependent on the diameter of the fractured bone and the patient s weight, but is chosen at the surgeon s discretion. Screw size is dependent on the requirements of the plate being applied as the screw holes accommodate a certain size. After choosing the appropriate plate and screw, a pilot hole is drilled for the chosen screw size using a drill bit and drill sleeve for accurate alignment. While conventional plating systems allow the surgeon to angulate the applied screws when inserted, studies have shown that insertion perpendicular to the plate provides the highest resistance to pullout in normal bone and forty degrees off-axis the lowest.[43] Different types of drill guides may be used to facilitate the intended function of the plating system. These include a universal guide to center the drill in the plate hole for a Neutrally applied screw for Neutralization, Bridging, and Buttress modes and an eccentric guide to offset the drill to enable Compression mode.[5] Figure depicts the difference in screw application using the two guides. When securing holes in an eccentrically drilled screw hole, the fragment ends of the fracture being secured glide toward each other and apply a compressive force at the fracture site. This is only possible in fractures that are transverse or minimally oblique, to ensure the fragment ends do not slide when exposed to compressive axial forces along the bone. 21

31 Figure 17: Eccentric Drill Guide (Left). Universal Drill guide (Right). The hole depth is measured and the correct screw length is used. If the correct length is unavailable, the next longer screw is chosen. If a standard screw is being applied, threads are first tapped in the hole and the screw is then inserted through the plate and bone and tightened with a torque-limiting screwdriver. This process is repeated for the remaining screw locations chosen by the surgeon to fit the plate to the bone. To ensure axial alignment of the plate to the bone, screws are first applied at the ends of the plate most distal to the fracture site. Screws are then applied to the holes most proximal to the fracture and then to the remaining holes chosen to secure the plate to the bone.[5] Alternatively, if the alignment is straightforward, the surgeon may choose to first fix the plate to the bone at the holes most proximal to the fracture gap, and then alternate sides moving toward the end of each plate. Regardless of the procedure used, screw torque is confirmed after all have been seated. These two securing methods are labeled as A and B in Figure 18. Figure 18: Screw Securing Orders "A" and "B". 22

32 2.4 Bone Loading There are four predominant mechanical testing methods used in the evaluation of bone and plating systems: tension, compression, bending, and torsion. These sufficiently calculate various mechanical properties of bone Tension Tensile testing is a simple method of determining the mechanical properties of both cortical and cancellous bone, though the specimen required is usually large. Prior to testing, specimens are modified to form a dog bone-like shape to ensure a uniform strain in central portion of the specimen. This test is not physiologically relevant as in vivo bone is not primarily loaded in tension, though it may still be used to calculate Young s modulus, ultimate strength, and yield strength.[6, 30] Only a limited number of studies evaluate the tensile properties of healing bone. In one study, bone healing was investigated in an osteotomy of the diaphysis of the right sheep metatarsal as a function of gap size and stability. Specimens containing callus were evaluated nine weeks after surgery and were found to have a decreasing tensile strength with greater initial interfragmentary gap.[44] For example, a one millimeter initial gap corresponded to a tensile strength eight times lower than normal bone; a six millimeter gap resulted in a tensile strength 36 times lower than normal bone Compression Compression testing is more commonly used to test cancellous bone as this is its most physiologically relevant loading mode. Compression tests are advantageous as they require smaller specimens compared to tensile tests, thus simplifying the mechanical testing. Most importantly, compression testing more closely simulates in vivo loading conditions, especially in vertebrae.[6] Similar to tensile testing, few studies have reported compressive properties of healing bone. The same study analyzing tensile properties of healing bone collected from a transverse osteotomy in the midshaft of sheep metatarsal nine weeks after surgery reported compressive testing results. Testing showed a great variation in mean indentation stiffness depending on callus location, gap size, and stability.[44] A similar technique was used in a study using the canine femoral osteotomy model. Specimens showed an increase in indentation stiffness from 2% to 25% from two to twelve weeks after surgery. The change in local stiffness over time also correlated significantly to the increase in maximum torque and torsional stiffness.[45] Bending Bending tests can be performed either in three-point or four-point loading modes. Threepoint bending involves a simply supported beam with one load applied between the supporting ends. Four-point bending applies two loads between the supporting ends. Bending tests allow tensile stresses to be present on one side of the specimen while the other side experiences 23

33 compressive stresses. As bone is usually weaker in tension than in compression, failure most commonly occurs on the tensile side of the specimen.[6] Bending is an essential test when evaluating bending strength and stiffness properties of long bones.[6, 7] Three-point bending is a common test, but causes a high shear stress near the middle section of the specimen. With the addition of a second, equally-spaced, load applied to the specimen, pure bending is achieved in the middle section of the specimen without the presence of transverse shear stresses. Stress, strain, and Young s modulus can be calculated as well, in addition to structural stiffness of the bone. A study conducted on healing metatarsal bone nine weeks after surgery evaluated its bending stiffness and showed that bending stiffness increased in specimen with smaller gap sizes and higher fixation stability.[6, 44] Torsion Torsion tests are an effective way of measuring biomechanical properties of bone in shear. When loaded, shear stress varies from zero at the center of the specimen, to the maximum at the surface.[6] Shear stress and shear modulus are properties which can be calculated from acquired torsion data.[46] Depending on the requirements of the test, structural strength, fatigue strength, and stiffness can be calculated. This is depicted in Figure 19(c),(d). A study conducted by White et al. defined four different stages of fracture in torsion during the process of bone healing. Using the rabbit tibia as a model, partial and full failure was found to occur through the original fracture site during low and high callus stiffness. Paavolainen et al. performed a study analyzing the mechanical properties of tibio-fibular bone fixed with a DCP plate. It was found that the torsional stiffness, strength, and toughness of healing bone peaks between six and nine weeks after surgery. Waris performed a similar study where a DCP plate was fixed to rabbit tibio-fibular bones. It was found that the torque moment at fracture, energy absorption, rigidity, and angular deformation increased from 3 to 12 weeks after surgery. 24

34 Figure 19: Bone Loading. (a) Axial tension/compression. (b) Bending. (c) Torsion. (d) Shear. Adapted from Athanasiou et al.[6] Plating System Load Distribution As a result of applying an internal fixation system to bone, additional forces and stresses interact with the system and are dependent on the design of the system and the loading mode. These load distributions are an important consideration when designing an implantable plating system. When subjecting bone to axial compression and bending loads, additional force interactions are present in the system. Figure 20 depicts the forces associated with applying a conventional plate, as observed by the additional forces at the plate/bone, plate/screw, and screw/bone interfaces due to the compression and friction forces created. The additional forces present in fractures secured with a locking plate system, as seen in Figure 21, are at the screw and bone interface as no contact or compression is required between the plate and the bone. Figure 20: Axial and Bending Load Distribution for Conventional Plating Systems.[5] Figure 21: Axial and Bending Load Distribution for Locked Plating Systems.[5] 25

35 Similarly, when subjecting the systems to torsional loads, Figure 22 and Figure 23 are adapted from Gautier and Sommer and depict the additional force interactions in these systems.[47] Figure 22: Torsion Load Distribution for Conventional Plating Systems. Adapted from [47]. Figure 23: Torsion Load Distribution for Locked Plating Systems. Adapted from [47] 2.5 Bridging the Gap Unlike medical devices for humans, no regulations exist in the Federal Food, Drug, and Cosmetic Act under the FDA requiring evidence of the safety and effectiveness of veterinary devices. Although several studies evaluate the functionality of one or two plating systems, no research has been conducted evaluating the mechanical properties of each plating system in a standardized fashion. Mechanical and clinical evaluation of plating system mechanics and bone healing effectiveness will allow for a better understanding of system behavior in vivo and will pronounce their benefits and faults. Evaluation of the eight most prevalent plating systems will help veterinary surgeons better treat their canine patients with the optimal device for the fractured bone by identifying plating characteristics which both support and inhibit fracture healing. 26

36 3. Goals The most prominent loading modes experienced by bone are bending and torsion. To simulate this loading, two types of mechanical tests will be performed: four-point bending and torsion. Mechanical tests will be conducted simulating a comminuted femoral fracture with a plating system bridged across the gap. To diminish any variation with regard to bone geometry, implant placement, implant shape, and gap size, a standardized model must be used. Through the normal gait of canines, numerous forces act upon their bones translating primarily to a combination of bending and torsional forces. Though the exact mode and intensity of loading is not known, it is important to analyze the response of the plating systems to these loading types.[48, 49] As described in ASTM F382, Standard Specification and Test Method for Metallic Bone Plates, cyclic loading should allow for one million cycles to be placed on the plating system without causing failure. Though it is important to analyze the cyclic loading properties in each loading mode, testing machine limitations will allow us to only conduct cyclic loading in torsion. The only previously conducted testing on these plating systems evaluated their single cycle to failure properties in torsion. In this evaluation, we will determine the various mechanical properties of these seven plating systems when simulating a comminuted fracture utilizing a synthetic bone model. 3.1 Specific Aim 1 Evaluate the mechanical properties of eight plating systems in single cycle to failure four-point bending. Utilizing ASTM F382 as guidance, four-point bending tests will be conducted to evaluate the mechanical properties of the various plating systems. The bending stiffnesses and strengths will be calculated for each construct and compared to determine if any significant differences exist. Furthermore, failure modes of each plating system will be noted. Due to limitations in the equipment used for the acute four-point bending tests, cyclic loading will not be performed at this time. 3.2 Specific Aim 2 Evaluate the fatigue strength of seven plating systems in cyclic torsional loading. A previous evaluation has identified the acute failure torque in these systems.[50] Utilizing the staircase method, samples will be cyclically torqued for evaluation. This will allow us to ultimately determine the fatigue strength of each plating system utilizing probability plots and Dixon-Mood analysis. Furthermore, failure modes of each plating system will be noted. 27

37 4. Experimentation Acute failure properties were evaluated for eight plating systems using four-point bending tests. Seven plating systems were evaluated for their fatigue strength using cyclic torsion tests. While limited biomechanical data exist for a few of these systems, no data exist evaluating them in a standardized manner. Samples were assembled in a fashion to simulate a comminuted fracture utilizing a synthetic bone model as the fractured bone fragments and plating systems to span the fracture gap. 4.1 Bone Model To remove variability encountered with cadaveric bone, a synthetic bone model was used in the mechanical testing. Short-fiber filled hollow epoxy (SFE) cylinders were previously validated to simulate the mechanical properties of bone.[51] This study compared the elastic modulus, maximum stress, and yield stress of synthetic bone models of various dimensions to published pit-bull and greyhound femur, humerus, and tibia results.[8] Based on their results, the bone model with a 20mm outer diameter and 3mm wall thickness best represented these mechanical properties of native bone. The model used is depicted in Figure 24. Figure 24: Bone Model Cross-Section. 4.2 Plates and Screws Eight different plating systems were evaluated when used as a bridging plate and tested in acute four-point bending, while only seven of these plating systems were evaluated in cyclic torsion. While both SS and TI LC-DCP systems were evaluated in four-point bending, only the SS LC-DCP system was evaluated in torsion due to limited sample quantities. These systems included two conventional, four locking plates, and one combination standard/locking hole plate. These plates were secured to the bone model with the appropriate screws corresponding to the manufacturer s specifications. These included standard and locking screws for both monocortical and bicortical purchase. Table 3 below summarizes the properties of the plates and screws used. It should be noted that only the locking holes were used in the combination hole LCP plate. In 28

38 addition, bicortical screws were used secure the ALPS plates at the hole most distal to the fracture gap on both bone model segments and monocortical screws were applied to the remaining holes. For all systems, stainless steel plates were mated to stainless steel screws while titanium plates were mated to titanium screws. Although, these plating systems differ in their screw hole spacing, screw size, and number of screws used to secure the plate to the bone, all available screw holes on the plate are used when assembling the samples. As this method is followed for all plating systems, it maintains uniformity in the sample assembly and allows for a comparison to be made between all plating systems without accounting for screw density. Screw density is the quotient formed by the number of screws inserted and the total number of plate holes, typically recommended to be below 0.5.[5] For plates smaller than twelve holes, screws should be applied to at least three holes per fragment. Table 3: Plate and screw system features. Plate Screw/Plate Size Plate Hole Screw type Cortex Penetration Plate/Screw Material DCP 3.5mm Conventional Standard Bicortical SS LC-DCP 3.5mm Conventional Standard Bicortical SS/Ti LCP 3.5mm Conv/Locking Std/Locking Bicortical SS SOP 3.5mm Locking Standard Bicortical SS Fixin 3.0/3.5mm Locking Locking Bicortical SS ALPS10 2.7mm Locking Locking Mono/bicortical Ti ALPS11 4.0mm Locking Locking Mono/bicortical Ti 4.3 Assembly tools Numerous tools were required to assemble all the samples. Two custom jigs were designed and manufactured to aid the veterinarians in uniformly assembling test samples. A drilling guide was fabricated to stabilize the bone models while aligning and fastening a plating system. A one-inch long SFE tube was used between the bone fragments to maintain a fixed distance between the bone models. The gap created is larger than that observed in previous tests, thus testing the construct more strenuously and mimicking a bridging osteosynthesis.[36, 52, 53] These are seen in Figure 25. A torque-limiting screwdriver, drill, and the required drill bits and bit sleeves were also used to aid in assembly. All samples were assembled by a veterinarian to ensure compliance with manufacturer and surgical specifications. 29

39 Figure 25: Drilling Guide and Spacer. A centering jig was developed to ensure bone models are properly centered when potting bone models for torsion samples. This is pictured in Figure Aim #1 Experimentation Figure 26: Centering Jig Sample assembly For each sample, two six-inch bone models were placed into the drilling guide, spaced using a one-inch SFE tube, and secured with the method described in While smaller gaps were utilized in previously conducted studies, a larger gap may better accentuate the differences between plating systems.[53] The spacer was then removed, thus simulating a comminuted fracture. Plates were centered on the bone models to ensure equal plate contact on each bone model. Furthermore, all screw holes on the plate in contact with the bone models were fastened. The DCP, LC-DCP, LCP, and Fixin plates were secured to the bone models using three bicortical screws in each bone model segment. The SOP plate was secured to the bone models using four bicortical screws in each bone model segment. The ALPS10 and ALPS11 plates were 30

40 secured with one bicortical screw on each bone model segment at the furthest position from the fracture gap. Three monocortical screws were used to secure the plate at the remaining screw holes on each bone model segment. All screws were loaded neutrally and a torque limiting screwdriver was used to ensure the insertion torque applied to each screw was standardized to 2.5 N-m Initial Mechanical Testing Initial testing was conducted using an Instron 5544 electro-mechanical testing machine. The loading anvils were placed distal to the screw most proximal to the fracture gap on each bone model segment. The bottom support anvils were placed at their furthest position allowed by the grip. The sample was preloaded to 5N and loaded at a rate of 0.1mm/sec until reaching the deflection limit of the machine. After conducting a single test, interference between the top loading anvils and plating system screws were noted to cause undesirable deflection and fracture of the sample s screws in the sample. Figure 27 depicts this deflection on a Fixin plating system. Furthermore, as screw spacing differs between plating systems, the constant adjustment of loading anvil placement would be detrimental to maintaining a controlled environment. Figure 27: Fixin Screw Deflection Final Testing Acute, single cycle to failure tests were conducted using the same Instron 5544 testing machine utilizing ASTM F382 as guidance.[7] The top loading anvils were spaced 14cm apart to be distal to the distal-most screw on the sample with the longest bridging plate. The bottom support anvils were placed 24cm apart, corresponding to the widest position allowed by the grip. Each sample was manually centered in the test fixtures, preloaded to 5N, and loaded at a rate of 0.1mm/sec until reaching the deflection limit of the machine. Load and displacement data were recorded at 10Hz. Four samples were tested for each plating system, totaling 32 samples. Figure 28 depicts a properly installed sample in the test fixture. 31

41 Figure 28: Properly Installed Four-Point Bending Test Sample Aim #1 Results After mechanically testing all the samples, the acquired data are evaluated to extrapolate the bending stiffness, bending structural stiffness, and bending strength. Bending stiffness is defined as the maximum slope of the linear-elastic portion of the load-displacement curve for each sample. Bending strength is defined as the bending moment needed to generate a 0.2% offset displacement in the sample. Finally, the bending structural stiffness is a normalized effective bending stiffness taking the test setup configuration into consideration. To calculate the various mechanical properties, anvil span measurements must first be noted. Loading span, h, is defined as the distance between one bottom support anvil and the nearest top loading anvil. Center span, a, is defined as the distance between the top loading anvils. Both span measurements are used in their millimeter form. These values correspond to 50mm and 140mm, respectively, according to our test setup. Figure 29 is adapted from ASTM F382 and depicts these measurement locations.[7] 32

42 Figure 29: Anvil Span Measurement Locations. Adapted from ASTM F382.[7] Graphs are first generated for each sample plotting load over displacement, a sample of which is seen in Figure 30. Before making any calculations, a linear regression, represented by segment OM, is performed on the region of the graph most closely representing linear-elastic deformation in the system. A line with slope equal to that of the regression is plotted at a 0.2% offset to the linear regression to ensure calculations are based on a failure load in the plastically deformed area of the plot. This is represented as segment BN, which intersects the loaddisplacement curve at point P, labeled as the proof load, and is the load after which plastic deformation occurs. Figure 30: Sample Load-Displacement Curve. 33

43 The slope of segment OM is equal to the bending stiffness of the plating system, represented as K in the equations below. Using Equation 1, the offset value, q, is first determined in order to draw segment BN and determine the proof load. Because grip setup remains the same throughout the testing process, the calculated offset value of 0.028mm is used in calculating proof load for all samples. Using Equation 2, bending strength is now calculated. The final calculated mechanical property, bending structural stiffness, is calculated to normalize the stiffness data based on test setup configuration. The evaluated plating systems have different geometries and screws required to secure the plate to the bone, however, the test configuration does not change, and therefore we expect the normalized stiffness calculation for each system to be of similar order to its bending stiffness. Equation 3 is used to calculated this with a, h, and K as inputs. Equation 1: Offset Displacement Calculation. Equation 2: Bending Strength Calculation. Equation 3: Bending Structural Stiffness Calculation. ( ) The calculated mechanical properties are then analyzed using a one-way ANOVA with Tukey pairwise comparison to compare the values between all plating systems. Statistical significance will be set at p< Aim #2 Experimentation The up-and-down, or staircase, method was used to evaluate each plating system. Staircase testing is used to evaluate the fatigue strength of a material at a specific fatigue life.[54] In our application, a starting load is first chosen and torqued to a specified number of cycles. If the sample fails before reaching the predetermined number of cycles, the sample is designated as a failure and a subsequent sample is tested at a load decreased by a constant value. If a sample does not fail it is designated as a runout and a subsequent sample is tested at a load increased by the same constant value. This process is repeated for all samples. To ensure accurate results, it is recommended to test at least fifteen samples.[55] Sample Assembly The assembly of the torsion samples was similar to that of the four-point testing samples, however, an additional step of potting one end of each bone model segment in fiberglass resin was added prior to application of the plating system. The jig depicted in Figure 26 above was first used to center the bone model segment prior to potting it in fiberglass resin. A two-inch tall 34

44 and two-inch square PVC tube was inserted into the jig along with a bone model segment over the centering nub. Two pins were then inserted perpendicularly through adjacent faces of the PVC tube and bone model segment to keep the bone model segment centered upon removal from the jig. When removed, this setup was placed on a sheet of cellophane wrap and the tube was filled with resin to facilitate ultimate stability while testing. After the resin fully cured, the respective plate was secured to the bone models with the method described in Section The length of the bone model segment varied between plates to allow a fixed distance between the distal-most screw to the fracture gap and the top of the potting fiberglass resin. Table 4 below lists the bone model segment lengths for all plating systems and Figure 31 depicts multiple views of a complete bone model segment. Table 4: Bone Segment Lengths Plating System Bone Segment Lengths (mm) DCP 123 LCP 122 Fixin 118 ALPS ALPS SOP 130 LC-DCP 123 (a) (b) Figure 31: Complete Torsion Bone Model Segment (a) Front View (b) Top View Initial Mechanical Testing Initial testing was conducted using a MTS 858 Mini-Bionix hydraulic testing machine, utilizing the sample assembly methods from the four-point bending tests described in Section and machine grips from previously conducted acute torsion testing. This setup did not 35

45 include potted bone model ends, and was secured in the grip using screws to provide stability through a friction fit. This setup is seen in Figure 32. Figure 32: Initial Torsion Test Setup. Initial testing was conducted cycling the samples at 1Hz torqued to 60% of the plating system s failure torque in both positive and negative directions, as determined through prior testing. The maximum torque value from each positive and negative cycle was recorded. It was found that the grips used to secure the constructs allowed slight movement in the construct when cyclically tested, thus leading to inaccurate acquisition of data. Figure 33 plots the maximum positive torque values and corresponding angles recorded during a portion of the test. Though the torsion testing was torque driven, slip within the grip did not allow the specified torque value to be consistently attained. In an effort to reduce any slip throughout the test, the ends of the sample were potted in fiberglass resin and new machine grips were manufactured. 36

46 Figure 33: Positive Cycle Data for ALPS11 Cycled at 2Hz Final Testing Samples for final testing were assembled using the jig described in Section A pair of grips were designed and manufactured to accommodate the new sample assemblies in the testing machine. These grips are made of aluminum and their geometry allows for a close fitting sample. Screws are tightened on each face of the jig to provide supplementary stabilization prior to testing. Figure 34 depicts a properly installed sample end in the jig. Figure 34: Assembled Bone Segment End. No standard currently defines a method to evaluate fatigue properties of plating systems in torsion. Based on previous research and accelerated methods of analyzing material fatigue properties, sixteen samples of each plating system were tested in torsion using the staircase method to determine their respective fatigue strength.[54, 56-58] The first sample of each plating 37